Cantilever sensors for molecule detection

ABSTRACT

A method of preparing a cantilever sensor for measuring biochemical interactions and their associated stress wherein a cantilever having two sides is coated on one side with at least a gold layer and both sides of the cantilever are functionalized with a self-assembled monolayer (SAM) of a probe molecule by incubating the cantilever in a solution having a concentration of the probe molecule of between 1 to 1000 μM. The unpassivated cantilever sensor comprising a layer coated on one side with a coating comprising gold and being unpassivated on the opposite side, wherein both surfaces comprises a self-assembled monolayer of a probe molecule in which the surface area occupied per probe molecule varies in the range 0.4-1.5 nm2, enabling the stress at the gold top surface that is not cancelled out by a counter stress from the bottom surface so that accurate quantitation of a target molecule is achieved.

RELATED APPLICATION

This application is a Divisional Application of U.S. patent applicationSer. No. 15/191,056, filed Jun. 23, 2016, which is herein incorporatedby reference in its entirety.

TECHNICAL FIELD

The present invention relates to cantilever sensors and cantileversensor arrays having enhanced selectivity and sensitivity for thedetection of molecules in body fluids.

BACKGROUND ART

Cantilever sensors have attracted considerable attention over the lastdecade because of their potential as a highly sensitive sensor platformfor high throughput and multiplexed detection of proteins, nucleic acidsand other molecules.

Biological specificity in detection is typically achieved byimmobilizing selective receptors or probe molecules on one side of thecantilever using surface functionalization processes.

Biomolecular interaction between the immobilized receptors or probes anda coupling molecule in a body fluid causes a measurable bending of thecantilever. This nanoscale deflection is caused by a variation in thecantilever surface stress due to the biomolecular interaction and can bemeasured by optical or electrical means, thereby reporting on thepresence of specific molecules and their quantitation in the body fluid.

The cantilever bending is a function of the number of molecules bound tothe probe molecules on its surface.

Biosensing technologies based on cantilever arrays have the potential ofsatisfying the need for multi-target detection with high sensitivity andselectivity using very small volumes of sample.

When cantilevers are made softer with very small force constants, theycan measure forces and stresses with extremely high sensitivity.

The very small force constant (typically less than 0.01 N/m) of acantilever allows detection of surface stress variation due to theadsorption (or specific surface-receptor interaction) of molecules.

Known cantilever sensors are typically made of rectangular silicon, orpolyamide polymer materials coated with a layer of gold on one side (thetop surface) and with non-reactive layer of molecules such as PEG-silaneon the other side (the bottom surface). The PEG-silane coating of thebottom surface is called passivation layer. The purpose of passivationof the bottom surface is to help avoiding unwanted functionalization ofthe bottom surface with receptors or probe molecules, consequentlypreventing probe molecule (ligand) adsorption that would alter sensingresults. Receptors or probe molecules are typically immobilized on thecantilever top gold surface using, for example, alkanethiol chemistry.

The need for mass-produced, miniature microcantilever arrays havingunprecedented sensitivity for label-free biodetection applications, suchas toxin, protein, drugs and antibody detection, DNA hybridization,selective detection of pathogens etc. is significant. However,improvements in cantilever sensitivity and selectivity are far fromfinished and there is still a long felt need in the sector of label-freemolecular detection.

Passivating of the underside of the cantilever to prevent unwantedligand adsorption is lengthy and often requires tedious optimization.For example, on average more than 60 minutes is necessary to passivate asingle micro-cantilever array with PEG-silane. Moreover, the detailedinvestigation of Si surface passivation shows that the process ofcantilever underside coating to prevent unwanted adsorptions is far fromcomplete and therefore needs further optimization.

Therefore, it would be desirable to provide cantilever sensors withoutthe need of a passivation layer but, at the same time, having the sameor even an enhanced selectivity and sensitivity with respect to theknown passivated cantilevers.

SUMMARY OF THE INVENTION

The present invention refers to a process for immobilizing probemolecules (such as receptors or antibodies) able to interact with aligand or drug molecule on nanomechanical cantilevers so that they canfunction, in a reliable, selective and sensitive way, without the needto passivate the underlying surface.

By tuning the probe molecule concentration (i.e. by working in aconcentration range of 1 to 1000 μM), the impact of the underlyingsurface SAM functionalization can be minimized.

Moreover, the area per probe molecule (i.e. the probe molecule spacing)on the cantilever surface influences coupling molecule/probe moleculebinding and can affect the mechanical stress and hence on thesensitivity and selectivity of the sensor.

The process of the disclosure therefore yields a cantilever having thesame or even an enhanced selectivity and sensitivity with respect to theknown passivated cantilevers without the drawbacks of the passivationprocedure (which is lengthy and often requires tedious optimization).

In another aspect, the disclosure provides a method for detectingmolecules in body fluids having improved selectivity and sensitivity.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1a-b show the functionalization of both surfaces of the cantileverwith a self-assembled monolayer (SAM) of vancomycin (Van) susceptiblereceptor (VSR ˜0.6 kDa) analogues of the bacterial cell wall precursorsthat present uncrosslinked peptide motifs terminating in the sequencelysine-D-alanine-D-alanine.

FIG. 2a show the differential cantilever bending signal for 1 μM, 50 μMand 1,000 μM VSR against Van (fixed at 250 μM) to investigate the effectof surface chemistry on stress signaling. The differential PEG referencesignal is shown in black. The cantilevers were found to bend downwardsdue to steric and electrostatic repulsive interactions between boundligand-receptor complexes.

FIG. 2b depicts the semi-logarithmic plot showing measured differentialsurface stress response as a function of [VSR] in solution against Van(fixed at 250 μM), superimposed on the results of the fit using equation(1) (solid line), derived from formula (1). Regime I for VSR representsthe initial stages of the self-assembly of molecules, and regime II hasa complete SAM of molecules on both top and bottom cantilever surfaces.

FIG. 2c shows the differential SPR response signals for 0.1 μM, 1 μM and100 μM VSR against Van (fixed at 250 μM) to investigate the effect ofreceptor concentration on signal amplification.

FIG. 2d is the semi-logarithmic plot showing the measured differentialSPR response signal as a function of [VSR] in solution against Vanconcentration (fixed at 250 μM), superimposed on the results of the fitusing a mathematical model simulating a cantilever in which there is nofree available Si bottom surface (not reported in the disclosure).

FIG. 3 is a plot showing the measured normalized surface coverage(symbols) obtained using X-ray photoelectron spectroscopy (XPS) as afunction of the molar fraction of receptor (diluted with PEG insolution).

FIG. 4a shows the differential cantilever bending response signals insodium phosphate buffer solution for defined percentage ratios of 30%,70%, 90% and 100% VSR (diluted with PEG in solution) fixed at a totalreceptor solution concentration of 1 μM, at which the net cantileverstress signal contribution from the underlying Si reactions isnegligible when exposed against Van at 250 μM. A negative signalcorresponds to a compressive surface stress, which results in acantilever downward-bending deflection.

FIG. 4b shows a schematic representation of the coupling approach.Receptor molecules (Y shape) covalently bind to Au (top) surface viathiol groups that are created through the modification of reactiveN-hydroxysuccinimide (NHS) ester groups with primary amines. SAMterminating with PEG or OMe were incorporated between the receptors in adefined ratio to enhance the biospecific binding efficiency of thesensing layer by reducing non-specific binding on the Au surface.Passivation of the cantilever Si surface was achieved using PEG-silaneto block a non-specific underside reaction.

DETAILED DESCRIPTION OF THE DISCLOSURE

The process of preparation of an unpassivated cantilever comprises thesteps of:

-   -   1) Providing a microcantilever sensors having two sides;    -   2) Coating one side of the cantilever with at least a gold        layer;    -   3) Functionalizing both sides of the cantilever with a        self-assembled monolayer (SAM) of a probe molecule by incubating        the cantilever in a solution having a concentration of the probe        molecule of between 1 to 1000 μM.

The cantilever sensor used in the present process is preferably includedin an array comprising at least eight cantilevers. In a preferredembodiment, the process of the disclosure is a process of preparation ofan array of unpassivated cantilever, preferably an array of at leasteight unpassivated cantilever.

Preferably, the cantilever of the disclosure has a rectangular shape.Typical sizes of the cantilever sensor are: 500 μm long, 100 μm wide and1 μm thick.

The coating of one side of the cantilever with at least a gold layerpreferably includes the deposition of a first (or base) titanium layer(called the adhesion layer) and then of a top gold layer.

The coating of step b) is achieved by using any method known in the art.Preferably, the Au coating is prepared by using any of the known

physical thermal vapor deposition (PVD) methods (for example, thethermal evaporation technique) or any of the known PVD techniques, suchas the electron beam evaporation technique.

Metal coating preparation includes deposition of a first titanium layerfollowed by deposition of a gold layer under vacuum until the desiredthickness is achieved. Typically, the thickness of the titanium layer isbetween 1 and 5 nm. The thickness of the gold layer is typically between5 and 30 nm, preferably between 10 and 20 nm.

Self-assembled monolayer (SAM) refers to organic molecule assembliesthat form spontaneously on surfaces (for example by adsorption) and areorganized into more or less large ordered domains. Typically, moleculesthat assemble into monolayer possess a head group that has a strongaffinity to the surface and anchors the molecule to it, a tail and anend functional group. Common head groups include thiols, silanes,phosphonates, etc.

The self-assembled monolayer of the disclosure preferably is analkanethiol self-assembled monolayer in which the alkanethiol moiety isthe linker between the probe molecule (i.e. functional group) and the Auand/or Si surface, as depicted below:

Preferably, the alkanethiol linker is an alkanethiol polyethylene glycolmoiety.

The linker interacts with the Au and/or Si surface through the terminal—SH residue and is covalently attached to the probe molecule through the—OH group of the polyethylene glycol moiety. The interaction between theAu and/or Si surface and the alkanethiol linker is a semi-covalent typeof interaction due to the strong affinity of sulfur for these metals.

In a preferred embodiment, the alkanethiol linker isHS(C₈₋₁₅)alkyl-(OCH₂CH₂)_(n)OH, wherein n=2-5.

Preferably, the alkanethiol linker is HS(OCH₂CH₂)₃OH, that binds to theAu and/or Si surface of the cantilever via the —SH residue and to theprobe molecule via the —OH group, as depicted below:

The probe molecule can be any molecule able to interact with specificityand sensitivity with another molecule (a coupling molecule), thusgenerating ligand-receptor or drug-receptor or sequence-specific DNA orRNA hybridization-type interactions or antibody-antigen interactions.

The probe molecule preferably is a receptor able to provideligand-receptor or drug-probe binding or a probe molecule able toselectively hybridize to a complementary DNA or RNA sequence or anantibody able to provide antibody-antigen interaction.

For example, the receptor is selected from: the vancomycin susceptiblereceptor (VSR), monoclonal human immunodeficiency virus antibody(anti-p24), factor (VIII) antibody (anti-Factor (VIII)), polyclonalanti-prostate-specific antibody (anti-PSA), or combinations thereof.

The coupling molecule is a ligand, a drug molecule, a protein, anantigen, a hybridizing nucleic acid sequence, or combinations thereof.For example, the coupling molecule is vancomycin, glycoprotein p24,factor (VIII) and prostate specific antigen (PSA).

The probe molecule is attached with the linker, preferably analkanethiol linker, prior to the preparation of the incubating solution.

A solution of the probe molecule bound to the linker is prepared bydiluting the molecule in an organic solvent solution (such as ethanol ormethanol solution) to yield a concentration of probe molecule of 1 to1000 μM, preferably 30 to 100 μM, more preferably 20 to 60 μM, even morepreferably about 50 μM. The cantilever is incubated in the solution for10-60 minutes, preferably from 10-30 minutes and in any case until a SAMis formed.

By working in the probe molecule concentration of 1 to 1000 μM, asurface area occupied on the Au side per probe molecule of 1.5 to 0.4nm², preferably of 0.5 to 0.6 nm² is obtained and approximately 4.7 to1.6 nm² in case of Si underside. Higher values of surface area for probemolecule (such as 1.6 or 0.6 nm²) are obtained when lower concentrationsof probe molecule are used, i.e. when the probe molecule concentrationis below 50 μM. Lower values of surface area per probe molecule (such as0.4-0.5 nm²) are obtained when higher concentrations of probe moleculeare used, i.e. when the probe molecule concentration is 50 μM. The areaper probe molecule achieves the highest values when the probe moleculeconcentration is between 30 and 100 μM, preferably between 20 to 60 μM,more preferably around about 50 μM.

By working in the above ranges of probe molecule concentrations, an areaper probe molecule on the cantilever Au side of 0.4 to 1.5 nm²,preferably of 0.5 to 0.6 nm², is achieved, which yields a compressivestress of the cantilever (in the presence of a coupling molecule to bedetected in a body fluid) of 40 to 60 mN m⁻¹. This range of compressivestress is approximately twice more sensitive than the one obtained fromthe known cantilever (about 33 mN m⁻¹).

The area per probe molecule is calculated using X-ray photoelectronspectroscopy (as described in the example section of this disclosure).

Working in the range of the probe molecule concentrations in solutiondescribed above allows minimizing the amount of SAM that is formed onthe Si bottom side of the cantilever consistent with their negligibleimpact on cantilever stress. Accordingly, the influence of the negativecontributions of the interaction between the SAM formed on the Si sideof the cantilever and the coupling molecule to the overall stress signalis minimized. Besides being a minimized influence, the negativecontribution to the stress signal can be taken into account in order tocalculate the net change in stress, which can thus be correlated in areliable manner to the concentration of the coupling molecule to bedetected in a body fluid.

The inventors of the present disclosure have devised a mathematicalformula (1) to represent the simultaneous interactions at the Au and Sisurfaces, where the surface stress is defined by coupling molecule/probemolecule complex interactions.

The net change in stress is expressed as:

$\begin{matrix}{{\Delta \; \sigma_{eq}} = {{{\sigma_{\max}({Au})}\left( \frac{\lbrack{Ligand}\rbrack^{n}}{{K_{d}^{n}({Au})} + \lbrack{Ligand}\rbrack^{n}} \right)} + {{\sigma_{\max}({Si})}\left( \frac{\lbrack{Ligand}\rbrack^{m}}{{K_{d}^{m}({Si})} + \lbrack{Ligand}\rbrack^{m}} \right)}}} & (1)\end{matrix}$

where the first and second terms quantify the stress changes at the Ausurface and Si surface, respectively. For the associated stress to causean effective cantilever downward bending (compressive) with inclusion ofthe bound complex, the Au top expands, meaning the underlying Si surfaceundergoes contraction (tensile). Alternatively, other models known toone skilled in the art may be used.

In formula (1), the constants σ_(max)(Au) and σ_(max)(Si) are themaximum stresses when all accessible binding sites on the surfaces arefully occupied.

K_(d)(Au) and K_(d)(Si) are the equilibrium dissociation constants forthe Au and Si surfaces, respectively, and n and m are the stoichiometriccoefficients of the reactions.

It has been found that the equation formula (1) is a good mathematicalmodel to calculate the net change in stress (and accordingly theconcentrations of the ligand/drug molecule that is to be detected) whenthe starting concentrations of the probe molecule in solution is betweenof 1 to 1000 μM, preferably 30 to 100 μM, more preferably 20 to 60 μM,even more preferably about 50 μM. Indeed, it has been found that thevalue of K_(d)(Au) correspond to the value of K_(d)(Au) obtained from amodel simulating a cantilever in which there is no free available Sibottom surface, only when working in the above probe molecule solutionconcentrations. Instead, it was found that K_(d)(Au) increases by morethan an order of magnitude when the probe molecule concentration isbelow 1 μM or greater than 1000 μM. The increase in K_(d)(Au) at lowprobe molecule concentration, is due to the small net cantilever stresssignal contribution from both surfaces, for the reason that a completemonolayer is not formed at this concentration. Conversely, when theprobe molecule concentration is greater than 1000 μM, a largecontribution from Si reactions comparable to that from the Au topsurface obtained, results in a reduction in the net stress signal.

Thus, when either the probe molecule concentration is too low or higherthan 1000 μM, the extracted K_(d)(Au) values are artifacts.Consequently, a sufficient equilibrium net stress signal Δσ_(eq) isdesired for accurate binding analysis.

These findings demonstrate that the unpassivated cantilever of thepresent disclosure possess superior performance in terms of detectionsensitivity and specificity only when the starting probe moleculeconcentration in solution is included in the above ranges.

It has also been found that stress generation efficiency (and thereforethe cantilever sensitivity and selectivity) is also influenced by theprobe molecule spacing on the cantilever surfaces. In particular, it hasbeen found that, depending on the size of the incoming ligand or drug,the spacing must be tuned to allow perfect matching. When the incomingligand to be detected is a small molecule, such as an antibiotic, then asmaller spacing between the probe molecules can be utilized to creategood matching between probe molecule/coupling molecule. In contrast, forlarge molecules such as proteins, the probe molecule spacing must belarge in order to enable matching to incoming coupling molecules. Thismeans that in order to get large stress signal useful for sensitive andspecific sensing, the probe molecule spacing may be adjusted to obtainperfect matching between a probe molecule and a coupling molecule. Thereis no standard probe molecule spacing that is universal for all couplingmolecules.

It has been found that when the coupling molecule to be detected is asmall molecule, the surface area occupied on the cantilever Au side perprobe molecule is typically in the range 0.4-1.5 nm², preferably 0.5-0.6nm² in order for the cantilever to possess good sensitivity andselectivity. When the molecule to be detected is a large molecule, suchas a protein, the surface area per probe molecule is typically in apreferred range 0.6-1.2 nm², in order for the cantilever perform atbest.

It is to be noted that these values of surface area per probe moleculeare consistent with the surface coverage values obtained when workingwith a probe molecule concentration in solution comprised between 1 to1000 μM, preferably 30 to 100 μM, more preferably 20 to 60 μM, even morepreferably about 50 μM (which allow minimizing the quantity of probemolecule on the cantilever Si surface).

The present disclosure refers also to an unpassivated microcantileversensor comprising a range of materials including silicon layer coated onone side (or surface) with a coating comprising Au and being uncoated orunpassivated on the opposite side (or surface). The Au coated surface isfurther coated with a self-assembled monolayer of a probe molecule,wherein the surface area occupied per probe molecule is in the range0.4-1.5 nm², preferably 0.5-0.6 nm² or 0.6-1.2 nm² according to the sizeof the molecule to be detected. The probe molecule is able to bind in aselective way to the molecule to be detected in a body fluid, therebyforming a complex that causes bending of the cantilever. Bending of thecantilever generates a stress response that can be detected andcorrelated to the molecule concentration in the body fluid.

The cantilever coating comprising Au can also comprise a first layer oftitanium (called adhesion layer) on top of which the Au layer isdeposited.

The thickness of the titanium layer is between 1 and 5 nm. The thicknessof the gold layer is between 5 and 50 nm, preferably between 10 and 20nm.

The cantilever sensor of the present disclosure is preferably includedin a sensor array comprising at least 8 unpassivated cantilevers,although less or longer more unpassivated cantilevers can be used.

The probe molecule and the self-assembled monolayer are as describedabove.

The cantilever of the disclosure is able to detect the presence of amolecule in a body fluid with femtomolar sensitivity.

The present disclosure refers also to a method for detecting thepresence of a ligand/drug molecule in an ex-vivo body fluid, such asblood, plasma, urine, saliva, sweat or sputum (or combinations thereof),comprising the steps of:

-   -   1) Providing an unpassivated cantilever sensor according to the        present disclosure;    -   2) Contacting the cantilever with a body fluid containing the        molecule to be detected;    -   3) Detecting the response signal due to cantilever bending;    -   4) Correlating the response signal to the presence or absence of        the molecule to be detected and, in case of presence, to the        concentration of the molecule in solution.

The unpassivated cantilever sensor can be an array of at least 8unpassivated cantilevers or more.

The contact between the cantilever and the body fluid of step 2) isperformed for a period of about 5-30 mins, although shorter or longertimes can be used.

During the contact, the probe molecule selectively binds to the moleculeto be detected via a receptor-ligand, receptor-drug, antibody-antigen orhybridizing sequence-type of interaction, thereby forming a complex thatcauses mechanical stress and consequently a cantilever bending responsethat can be detected.

Detection of cantilever bending response in case of optical readout isperformed, for example, by using serial time multiplexed optical beammethod with a single position sensitive detector, although otherreadouts such as electronic or diffraction can be used. The laser spot(about 100 μm diameter) is aligned onto the free end of each sensorwhere the accuracy of alignment is confirmed by heating test. Theexpected precision of laser spot alignment is determined by calculatingthe bending variation at the maximum bending signals between individualcantilever sensors. Well aligned cantilevers at the maximum bendingsignals should yield a relative standard deviation of the bendingsignals of about ≤10%, preferably about ≤5%, between them. Correlatingthe response signal to the presence or absence of the molecule to bedetected includes a first step of calculating the net change in stress,using the mathematical model reported above, and then a second step ofassociating the net change in stress value to the presence of absence ofthe molecule and, in case of presence, to the concentration level of themolecule.

The molecule is considered not present in the body fluid when themeasured differential stress is equal to about zero, which correspondsto mechanical bending of the cantilever of about zero. As understood byone skilled in the art, the absence of a substance from a solution meansthat the substance concentration is below the sensitivity of theanalysis method. For the present method, a molecule to be detected isconsidered absent from a body fluid when the concentration of such amolecule is below the current limits of detection in femtomolarquantity.

A molecule is considered present in a body fluid when the compressivestress net signal is ≥0.02 mN m⁻¹.

EXAMPLES

Cantilever Metal Coating

Cantilever chips fabricated from Si (100) by IBM Research Laboratory,RUschlikon, Switzerland were first cleaned with freshly prepared piranhasolution (ratio 1:1 H₂SO₄ and H₂O₂) for 20 min. Arrays were thenthoroughly rinsed in deionized water before immersing in the secondfreshly prepared piranha solution for another 20 min, and again rinsedthoroughly with deionized water. Finally, the arrays were rinsed withpure ethanol and dried on a hotplate at 70° C. for 20 seconds. They werethen inspected using the optical microscope to confirm their cleanlinessbefore transferred to the evaporation chamber (BOC Edwards Auto 500,U.K.) for an overnight pumping. One side of the silicon cantileversurface was metalized using an electron beam evaporation with a 2 nm Tiadhesion layer followed by 20 nm of Au at a base pressure of ˜3×10⁻⁷mbar, and at an evaporation rate of 0.02 nm/s for Ti and Au,respectively, as measured directly above the source by a quartz crystalmonitor. Once the desired thicknesses were attained, the cantileverchips were left in the chamber for 2-3 hours as described previously tocool under vacuum before removing.

Functionalization of Cantilever Arrays

Case I: Simultaneous functionalization of Au(top) and Si(bottom)surfaces of cantilever arrays: To systematically alter both Au(top) andSi(bottom) surfaces into chemically active sensing layers, the arrays ofeight rectangular cantilevers were subjected to surface capturemolecules in micro-capillary glass tubes (King Precision Glass,Claremont, Calif., USA), arranged according to the cantilever pitch sizeof 250 μm. The alkanethiols of self-assembled monolayers (SAMs), namelyvancomycin susceptible receptor (VSR˜0.6 kDa) analogues to bacterialcell wall precursors that present uncrosslinked peptide motifsterminating in the sequence lysine-D-alanine-D-alanine and ‘inert’ SAMmolecule terminating in triethylene glycol (PEG) whose detailedstructural sequences have been described previously were diluted inethanol solution to yield total concentrations of: 0.1 μM, 1 μM, 10 μM,50 μM, 100 μM, 1000 μM, 2000 μM, 3000 μM and 4000 μM. Individual VSR andPEG concentrations were injected into micro-capillary glass tubes andcantilevers were then incubated inside the capillaries for 20 min,washed three times with ethanol and stored in distilled water untilusage. The functionalization procedure was carried out withoutpre-adsorbing resistive protein monolayer bovine serum albumin (BSA) orPEG-silane, known for blocking nonspecific interactions.

PEG was used as a cantilever reference coating material because of itsknown properties of minimizing or blocking biomolecular adsorptions(proteins/drugs) on the surfaces.

To tune the receptor footprint (area per each receptor molecule) forenhanced biodetectability and to achieve satisfactory quantitativedrug-target measurements, certain steps within the functionalizationprotocol were addressed. The first task was to define the percentageratio of VSR by incorporating a receptor molecule with secondSAM-forming molecule PEG. The ratio of VSR and PEG molecules dissolvedin ethanol was varied to yield a total concentration in the solution of1 μM, for which full surface coverage was established. The correspondingmixed ethanolic thiol solutions of VSR and ‘inert’ PEG mixed in theratio of 30%, 70%, 90%, and 100% were used to coat cantilever arrays.

Functionalization of SPR Sensor Chips with VSR Receptor Molecules

Case II: functionalization of SPR sensor chips with VSR receptormolecules: The plain Au-coated SPR sensor chips were covered with 100 μlethanolic thiol solutions of VSR and incubated for 20 min tosystematically alter a Au-coated surface of the SPR sensor chip intochemically active sensing layer before washing three times withultrapure ethanol. Binding and affinity experiments were performed onT100 BIAcore

SPR Instrument using the same solutions as in the cantilever experimentsfor the drug-target binding interactions.

Cantilever Si Underside Passivation

Case III cantilever Si underside passivation: The passivation of Si(bottom surface of cantilever) arrays of the eight rectangularcantilever probes was performed using2-methoxypoly(ethyleneoxy)propyl]trimethoxysilane (7 ethylene glycolunits, ABCR, Karlsruhe, Germany) or PEG-silane. This procedure wasachieved by incubating freshly functionalized cantilever sensor arraysin a mixture of 3 ml dry ethanol and 15 μl of PEG-silane for 30 min tocreate a PEG-silane layer on bare silicon under side surface. Next, thepercentage ratios of receptors (2B4F and anti-factor (VIII)) was definedand the surface molecular footprints were tuned. A surface linkermolecule, (HS-C11-Eg)₃-OCH₂—COONHS) (ProChimia Surfaces, Poland) whereEg is the ethylene glycol group and NHS is the N-hydroxysuccinimidegroup, was incorporated with a second SAM-forming molecule PEG(HS-C11-Eg)₃-OMe) where Me is the methyl group. The NHS and PEG werediluted in ethanol to yield the percentage ratios of 20%, 60% 80%, 90%,and 100% where the total concentration of the SAMs in the solution wasfixed at 2 mM.

The mixed ethanolic thiol solutions of NHS and PEG were subsequentlyused for the functionalization of cantilever arrays. The freshly coatedcantilever arrays with a mixture of NHS and PEG were incubated in sodiumacetate buffer solution (5 mM, pH 5.4) for 5 mins to activate thesurface for the coupling reaction. The solutions (50-100 μg/ml in sodiumphosphate buffer at pH 7.4) of variable domains of 2B4F (˜15 kDa)derived from VHH of llama single heavy chain antibodies and anti-Factor(VIII) (˜280 kDa) were injected across cantilever arrays and incubatedovernight at 4C for coupling reactions to occur at the activated NHSthiolated binding sites. After an overnight incubation, the microarrayswere subjected to a ‘capping’ procedure, using 1 M ethanolamine, pH 8.5to de-activate unreacted NHS thiols and rinsed thoroughly in PBS bufferat pH 7.4 three times and stored in sodium phosphate buffer at pH 7.4until usage.

SPR Functionalization with VHH 2B4F Using Carboxymethyl Dextran Matrix(CM5) and Plain Au-Coated SPR Sensor Chips

Case IV: SPR sensor chip functionalization with 2B4F antibody: PlainAu-coated SPR sensor chips were coated with 2B4F via NHS thiolatedsurface linkers. The 2B4F was immobilized on a plain Au-coated SPRsensor chip at 25° C. The procedures involved offline functionalizationand online in-situ functionalization to investigate their performance.For the online functionalization, the surface of the Au-coated SPR chipwas injected with the mixture of NHS and PEG in the ratio of 9:1 at 2 mMtotal concentration of both SAMs for 20 min at a flow rate of 5 μlmin⁻¹. The freshly functionalized SPR sensor chip was activated for 7min at a flow rate of 5 μlmin⁻¹ using a 5 mM solution of sodium acetatebuffer before the injection of 2B4F diluted in 100 mM sodium phosphatebuffer at pH 7.4 to a concentration of 100 μg/ml at a flow rate of 5 μlmin⁻¹ for 10 mins. The unreacted activated NHS groups were then blockedby using a 7 mins injection of 1 M ethanolamine, pH 8.5, at a flow rateof 5 μl min⁻¹. For the offline functionalization, the plain Au-coatedSPR chips were incubated with 2B4F solution for 20 min to systematicallyalter a Au-coated surface of the SPR sensor chip into a chemicallyactive sensing layer before washing three times with ultrapure ethanol.In case of the carboxymethyl dextran matrix (CM5), SPR sensor chip, anonline in-situ functionalization was used for all the experiments.

Solutions

Vancomycin and protein solutions: PBS: 0.1 M mono- and di-basic sodiumphosphate salts (Sigma-Aldrich, UK) were dissolved in ultrapure water(18.2 MΩ cm 10 resistivity, Millipore Co., Billerica, Mass., U.S.A.) andmixed to yield a pH value of 7.4. The buffer solutions were filteredusing 0.2 μM filter (Millipore) and subsequently used to dissolvegp140CN54 (subtype B/C) and gp140UG37 (subtype A). Differentconcentrations of these HIV-1 isolates were serially diluted from thestock (532 μg/ml).

Characterisation of VSR Using X-Ray Photoelectron Spectroscopy

The surface coverage of VSR was quantified using X-ray photoelectronspectroscopy (XPS), based on the previously reported method. In order toestimate the area of VSR ligand-presenting molecule on the goldsubstrate, the extensive research that has been done on simplealkanethiol monolayers on gold surface was utilized. For instance, it iswell understood that an ideally packed alkanethiol SAM on Au surfaceyields a packing density of ˜0.26 nm⁻². It is also known fromexperimental values that a PEG monolayer has a grafting density of ˜0.27nm⁻² and therefore by comparing the ratio of the S 2p/Au 4f signal ofthe VSR and PEG reference samples, the surface area per ligandpresenting molecule (Table S1) could be calculated.

TABLE S1 The surface area occupied per VSR molecule calculated fromX-ray photoelectron spectroscopy (XPS) data. VSR concentration Surfacearea per per VSR (μM) Molecule (nm²) 0.1 4.65 1 1.59 10 0.65 50 0.54 1000.55 1000 0.47 2000 0.44

The samples containing the lowest concentration of VSR gave the leastpacking density, with an area of ˜4.65 nm² per VSR molecule, whichcorrespond to a widely dispersed VSR layers. Whereas for highlyconcentrated samples, it was found that they gave denser packingdensities with an area of ˜0.44 nm² per VSR molecule (Table S1). Incontrast, for the lower concentrations, the grafting density was foundto vary across the range, yielding a dramatic increase in the area perreceptor molecule of ˜0.54 nm² per VSR molecule at 50 μM concentration.Additionally, as the VSR concentration was increased further, thepacking density was shown to increase rapidly (Table S1).

Finally, to determine which structural features of the surface thatregulate the access and insertion of ligand molecules to enhance theefficiency of stress transduction on the cantilever arrays, a mixture ofVSR/PEG solutions was prepared with a VSR molar fraction of 0.0, 0.05,0.1, 0.3, 0.5, 0.7, 0.9 and 1.0 (the ratio of 0.0 indicates pure PEGsolution and the ratio of 1.0 represents pure VSR solution). Highresolution scans of the N (1s), S (2p) and Au (4f) peaks were recordedusing an analyzer pass energy of 20 eV and by referring to the graftingdensity of 0.27 nm⁻² per PEG monolayer, the surface area occupied byeach VSR molecule in the mixed ratios was calculated. The results aresummarized (Table S2).

TABLE S2 The surface area occupied per VSR molecule (diluted with PEG insolution) calculated from X-ray photoelectron spectroscopy (XPS) data.VSR mole fraction Surface area per per VSR (μM) Molecule (nm²) 0.0 ∞ 0.34.54 0.5 2.20 0.7 1.20 0.0 0.64 1.0 0.44

Results

FIG. 1a-b show the functionalization of both surfaces of the cantileverwith a self-assembled monolayer (SAM) of vancomycin (Van) susceptiblereceptor (VSR ˜0.6 kDa) analogues of the bacterial cell wall precursorsthat present uncrosslinked peptide motifs terminating in the sequencelysine-D-alanine-D-alanine. To eliminate the artifacts that producenon-specific signals differential measurements were performed in whichreference polyethylene glycol(PEG)-coated cantilever bending signalswere subtracted from the receptor signals.

This functionalization was performed without pre-adsorbing a resistiveprotein monolayer of bovine serum albumin (BSA) or PEG-silane, which areknown to block non-specific interactions. Van was used as a reportermolecule because it reacts specifically with VSR to generate stress,which leads to cantilever bending deflections.

To probe the stresses due to antibiotic binding to different sensingsurfaces, Van was injected onto unpassivated cantilevers functionalizedwith VSR. The outcome after adding 250 μM Van is summarized in FIG. 2a .The bending response (shown in FIG. 2a ) is caused by interactions ofVan molecules at the surface, with the formation of a Van-VSR complexthat induces a local strain in the cantilever as well as carrying anelectrostatic positive charge under a physiologically relevantenvironment. The electrostatic repulsive and steric interactions betweenVan-VSR complexes create a compressive stress at the Au surface, causingthe cantilevers to bend downwards. The reference PEG-coated cantilevers,as expected, show no bending response against Van. The mechanicalresponse generated by the Van-VSR complex interactions increases withincreasing VSR concentration, but decreases at the high concentration of1,000 μM.

The effect of VSR concentration on signal amplification is summarized inFIG. 2b . The stress response to VSR concentration may be categorizedinto two regimes (I and II). Regime I represents the initial stages ofself-assembly of the molecules and is characterized by a sharp rise incompressive stress up to 52 mN m⁻¹, when the concentration of VSR insolution is 50 μM. This is approximately two times more sensitive thanprevious measurements, where the net stress was measured to be 33 mNm⁻¹. However, as the VSR concentration increases beyond 50 μM (regimeII), a significant decrease in stress signals down to σ_(max)≈5 mN m⁻¹was found. In contrast, zero differential stress was observed for thereference PEG-coated cantilevers. As a further measurement controluncoated Au and Si surfaces were used. The undetectable mechanicalresponse in the presence of Van is additional verification that theobserved deflection signal is caused by the interactions of Van withVSR. In general, the non-monotonic stress signal changes observed inFIG. 2b are not surprising given that the reported SAM formation on Sican give rise to negative contributions to the net cantilever stresssignal.

Ligand Sensing Based on Surface Plasmon Resonance

With stress signal reduction at the Au surface occurring at high VSRconcentrations (FIG. 2b ), the next objective was to confirm that thesechanges were caused by the opposing Si reactions. Therefore, acommercially available surface plasmon resonance (SPR) method where thedetection of biochemicals is at a single planar metal surface was used.

SPR detection is based on monitoring changes in the dielectricproperties caused by ligand adsorption. Accordingly, a series of bindinganalyses for VSR was performed, with [Van] kept constant at 250 μM tomatch the experimental conditions for the cantilever-based measurements.FIG. 2c shows that the differential SPR signal response increases withincreasing VSR concentration.

The SPR response features an S-shaped curve, with a steep rise, then aplateau when the receptor concentration increases beyond 10 μM (FIG. 2d). The SPR analysis of the signal response versus VSR concentrationshown in FIG. 2d remains constant even when [VSR] is extended to 1,000μM and, when compared with direct mechanical quantitation (FIG. 2b ),demonstrates that the reduction of stress signals at higher VSRconcentration is linked to Si reactions on the cantilever underside.These measurements provide the first demonstration that the directfunctionalization of cantilevers, without underside passivation (FIG. 1a), can be achieved by the effective tuning of receptor concentrations insolution described here. Previous measurements using cantilevers havefocused on one side only, but this analysis shows how the underlying Sisurface affects the overall mechanical response.

Effect of Receptor Surface Footprint on Ligand Binding

Although it is understood that receptor-ligand interactions in solutionare linked to stress generation, it is unclear how the surface footprintcorrelates with the concentration of receptors in solution. The resultsshowing that surface coverage is a function of receptor concentration insolution are summarized in Tables 1 and 2 below and FIG. 3. To study theimpact of receptor spacing on stress generation efficiency, a receptormolecule with a second SAM-forming molecule (PEG) was incorporated onthe Au surface without underside passivation.

PEG was chosen because it resists the unwanted adsorption of ligands byacting as a protein ‘repellent’. Moreover, it acts as a ‘spacer’ byvarying the distribution of receptors on the surface whilesimultaneously controlling the accessibility of ligands.

FIG. 4 shows the outcome when cantilevers were exposed to a constantantibiotic concentration at 250 μM Van with a defined ratio of VSR/PEG,where the total receptor concentration was fixed at 1 μM to minimize thenegative impact of Si reactions. With sparsely distributed receptor(˜30%) the cantilever deflection signal is negligible. However, when thereceptor concentration is increased to 100%, the surface packing densityis maximized and yields the highest stress (FIG. 4). These actions showthat the number of ligand-receptor interactions increases with coverage,although there is a threshold in the surface footprint to generate amechanical response, in good agreement with previous studies.

To examine the impact of coverage on signal amplification and to excludeany possibilities of the contributions from Si reactions, the cantileverunderside was passivated. To provide insight into the dependence ofstress generation on molecular size, testing was performed on theN-terminal fragment (VHH) of llama single chain antibodies 36 that havea molecular weight of about 15 kDa, some 25× larger than VSR but only10% of a conventional immunoglobulin. VHH are stable over a broadtemperature range (−80° C. to 80° C.) and are inexpensive to manufacturewith excellent expression yields from bacteriophage libraries. VHHraised against the human immuno-deficiency virus (HIV-1) trimericenvelope glycoprotein (gp140), as previously described, was chosen. OneVHH, termed 2B4F, was chosen because of its high specificity andsensitivity of binding to the gp140 by SPR. FIG. 4b shows the outcomeafter exposure to recombinant antigens derived from HIV-1 subtype A(gp140UG37, 140 kDa) fixed at 50 μM against a defined percentage ratioof 2B4F/PEG, where the total concentration was fixed at 2 mM.

The observed noise was probably caused by laser light scattering by theproteins. The response signal was not detectable at 20% relativeconcentration, but increased as the concentration was increased between80 to 90%. Surprisingly, 100% receptor concentration in solution wasfound to yield insignificant stress signals. Generally, these findingsreveal that the efficiency of stress generation for proteins is stronglydependent on the surface molecular footprint. In contrast, for smallmolecules such as Van (˜1.4 kDa), the stress is maximized when thereceptor packing densities is highest.

We claim:
 1. An unpassivated cantilever sensor having two sidescomprising a polyamide or silicon layer coated on one side with acoating comprising at least a gold layer and being uncoated orunpassivated on the opposite side, wherein the gold layer-coated surfacecomprises a self-assembled monolayer (SAM) of a probe molecule bound toa linker and wherein the surface area occupied per probe molecule rangesfrom about 0.4 to about 1.5 nm².
 2. The unpassivated cantilever of claim1, wherein the coating comprising at least a gold layer is a coatingcomprising a base titanium layer and a top gold layer.
 3. Theunpassivated cantilever of claim 2, wherein the thickness of thetitanium layer is between 1 and 5 nm.
 4. The unpassivated cantilever ofclaim 2, wherein the thickness of the gold layer is between about 5 andabout 50 nm.
 5. The unpassivated cantilever of claim 2, wherein thethickness of the gold layer is between about 10 and about 20 nm.
 6. Theunpassivated cantilever of claim 1, wherein the self-assembled monolayeris an alkanethiol self-assembled monolayer in which the alkanethiolmoiety is the linker between the probe molecule and the gold and/or anyother material including silicon surface of the cantilever.
 7. Theunpassivated cantilever of claim 6, wherein the alkanethiol linker is analkanethiol polyethylene glycol linker.
 8. The unpassivated cantileverof claim 6, wherein the alkanethiol linker isHS(C₈₋₁₅)alkyl-(OCH₂CH₂)_(n)OH, wherein n=2-5.
 9. The unpassivatedcantilever of claim 6, wherein the alkanethiol linker isHS(C₈₋₁₅)alkyl-(OCH₂CH₂)₃OH, that binds to the gold and/or any othermaterial including silicon surface of the cantilever via the —SH residueand binds to the probe molecule via an —OH group on the probe molecule.10. The unpassivated cantilever of claim 1, wherein the probe moleculeis a molecule able to interact with specificity and sensitivity with acoupling molecule to generate ligand-receptor interactions,drug-receptor interactions, antibody-antigen interactions,sequence-specific DNA interactions, RNA hybridization-type interactions,or combinations thereof.
 11. The unpassivated cantilever of claim 1,wherein the probe molecule is a receptor able to provide ligand-receptoror drug-receptor binding; a probe molecule able to selectively hybridizeto a complementary DNA or RNA sequence; or an antibody able to provideantibody-antigen interactions.
 12. The unpassivated cantilever of claim10, wherein the probe molecule is a vancomycin susceptible receptor(VSR), a monoclonal human immunodeficiency virus antibody (anti-p24), ablood clotting factor (VIII) antibody (anti-Factor(VIII)), or apolyclonal anti-prostate-specific antibody (anti-PSA).
 13. Theunpassivated cantilever of claim 1, wherein the surface area occupiedper probe molecule ranges from about 0.5 to about 0.6 nm².
 14. Theunpassivated cantilever of claim 1, wherein the surface area occupiedper probe molecule ranges from about 0.6 to about 1.2 nm².
 15. An arraycomprising at least 8 unpassivated cantilever sensors, each unpassivatedcantilever sensor according to claim
 1. 16. A method for detecting thepresence of a coupling molecule in an isolated body fluid, comprisingthe steps of: 1) providing an unpassivated cantilever sensor accordingto claim 1; 2) contacting the cantilever sensor with an isolated bodyfluid containing the coupling molecule to be detected; 3) detecting theresponse signal due to cantilever bending; and 4) correlating theresponse signal to the presence or absence of the coupling molecule tobe detected.
 17. The method according to claim 16, wherein the bodyfluid is blood, plasma, saliva, sputum, or urine.
 18. The methodaccording to claim 16, wherein the coupling molecule is a ligand, a drugmolecule, an antigen or a hybridizing nucleic acid sequence.
 19. Themethod of claim 18, wherein the coupling molecules is vancomycin,glycoprotein p24, factor (VIII), or prostate specific antigen (PSA). 20.The method of claim 16, wherein the contacting step is performed for aperiod of 5-30 minutes.
 21. The method of claim 16, further comprisingthe step of correlating the response signal to the concentration of thecoupling molecule in solution.